1. Field of the Invention
The present invention is directed toward a system and apparatus for accurately and rapidly measuring acoustic power flow at various points within the human ear canal.
2. Description of Related Art
In the field of audiology the chief objective of most diagnostic tools is to obtain an accurate measurement of wide band, middle ear power flow as estimated from the ear canal power flow. Power flow per unit area is known as acoustic intensity. An accurate reading of acoustic intensity is necessary for improved clinical diagnosis with respect to pathologies of the human auditory system. Accuracy alone, however, is insufficient to provide a clinically acceptable diagnostic system. The measurements must be obtainable rapidly and in a cost-efficient manner. Thus, the goal of the present invention is to provide a cost efficient, rapid, and accurate system of obtaining an acoustic intensity measurement within the human auditory system. Such a system will provide, inter alia, significant benefits in hearing screening programs leading to earlier discovery of ear pathologies.
Diagnostic tools such as air conduction audiometry, bone conduction audiometry, evoked response audiometry, and evoked otoacoustic emissions are critically dependent on an accurate measurement of the acoustic sound field. More specifically, acoustic intensity is a desirable measurement. Acoustic intensity is a measurement of the power flow per unit area. Unfortunately, measuring intensity directly is extremely difficult.
A more easily obtainable measurement is that of sound pressure as opposed to acoustic intensity. Pressure instruments are widely used in audiology. Pressure, however, only yields an intensity estimate when no standing waves are present, i.e., when power is flowing in one direction only, namely in sound fields that do not have acoustic reflection components. In sound fields having acoustic reflection components, standing waves will be produced and a pressure measurement is ineffective.
When reflections are not present, power flow is proportional to the square of the pressure multiplied by the area along the ear canal. In a calibrated pressure field having no reflection, power can be independent of frequency. However, when reflections are present in the ear canal then canal impedance is a finction of location in the ear canal, even if the pressure field is calibrated. In this case the square of the pressure does not characterize power flow. Sound pressure and acoustic intensity have a complex relationship when reflections are present in the ear canal. Thus, a single pressure measurement cannot determine the acoustic power flow in the ear canal unless the source transducer has been characterized. The source transducer is a small loud speaker/receiver combination placed in the ear canal. At least two independent measurements are necessary to do this. Characterization requires determining the source transducer's open-circuit pressure and its source impedance. Use of pressure as a measure of signal strength does not, however, take into consideration the source transducer open-circuit pressure or the source transducer impedance and thus cannot determine actual acoustic intensity in an ear canal having standing waves due to reflection.
In the early days of audiometry (circa 1930), sound levels were calibrated in a free-field (i.e. a field free of standing waves) where sound pressure and acoustic intensity are equivalent. Supra-aural headphones soon became popular because of their increased ease of use, acoustic isolation, and reduced low frequency calibration variability. These headphones were typically calibrated with a standardized acoustic coupler (i.e., artificial ear). This method of calibration improperly assumed that the acoustic impedance of the coupler was essentially the same as that of the ear being tested. As a result, considerable difficulty was encountered in developing a practical, yet reasonable way of accurately specifying hearing loss. One practical problem that emerged in defining the "normal threshold of hearing" was that different values for the auditory threshold were obtained on the same subject using different audiometric headphones, such as the American TDH-39 and the British STL headphones, although both headphones were calibrated on the same standard coupler.
A practical compromise was reached in 1969 with the standardization of the International Standard Reference Zero for audiometers. This standard is now universally accepted. In order to measure hearing level it is necessary to used a standard headphone calibrated in a standard coupler. A correction factor is needed if a transducer or coupler other than the standard headphone-coupler pair is used.
The above approach to the measurement of hearing level represents a practical compromise that works only moderately well below 4 kHz for normal adult ears. The approach remains, however, cumbersome and prone to error. Two important sources of error include inter-subject variability in the acoustic impedance of the tympanic membrane and in the cross-sectional area of the ear canal, and standing waves in the ear canal which result from an impedance mismatch between the tympanic membrane impedance and the ear canal characteristic impedance. These errors are sufficiently large above 4 kHz so as to render this method unusable.
The above sources of error are of particular concern when measuring hearing levels in infants and children because of the smaller physical size of their ear canals and the difference in acoustic impedance from that of the average normal adult ear. Similarly, the measurement of hearing levels in the presence of middle-ear pathology can also lead to error because acoustic impedance of the middle-ear is likely to deviate substantially from that of a normal ear. These two sources of error are often compounded in pediatric audiology because of a high incidence of otitis media in children.
A related problem is measuring acoustic signal levels produced by a hearing aid. The presence of an ear mold or an in-the-canal hearing aid results in a substantial change to the sound field within the ear canal. Under these conditions the predicted ear canal sound pressure level measured by a coupler can be very misleading.
Given the limits of current technology, one approach to this problem has been to increase the accuracy of the determination of the pressure not the intensity in the ear canal. For example, within the past ten years "real ear" measurement systems have become popular. These systems estimate the pressure in the ear canal in an attempt to lessen the uncertainty between a standard coupler and the ear being tested. When standing waves are present, however, pressure in the ear canal away from the tympanic membrane is not the same as the tympanic membrane pressure, nor does it characterize the acoustic power flow in the ear canal (i.e. the true intensity).
In the last five to ten years, evoked otoacoustic emissions, such as distortion products, have proven to be an important new method for characterizing the outer hair cell finction in the cochlear. Distortion product evoked otoacoustic emissions are small nonlinear cochlear retrograde signals. While this nonlinear measurement represents an important positive step in diagnosing hearing loss, it is also affected by standing waves due to middle ear reflections. When a calibration microphone is in a primary pressure null, created by a reflected (retrograde) pressure wave that partially cancels the forward traveling wave, the pressure at the measurement point and at the ear drum can differ by an arbitrarily large amount. Recently, 20 dB standing waves in adult ears sealed with an insertion--transducer have been observed at frequencies as low as 3.5 kHz. Under these conditions the ear canal acoustic intensity is not properly calibrated. Since distortion product evoked otoacoustic emission and other clinical measures depend on stimulus intensity the reliability of the calibration is critical.
This problem is exacerbated in cases of neonates and infants due to vemex in the canal within the first few days after birth, and middle-ear infections in infants and young children.
In all of these cases a significant percentage of the acoustic energy may be reflected by the middle-ear due to pathological middle-ear impedance mismatch.
The practical consequences of this impedance mismatch problem are substantial. Consider, for instance, a universal hearing screening program for infants. For every 1000 infants screened, we might expect only two or three to be cochlear-impaired (i.e. 0.2-0.3% of the population), and 50 to 100 (5-10%) to have some sort of middle-ear pathology (usually temporary). Both of these groups will test positive in an evoked otoacoustic emissions screening program. The middle-ear "positives" represent a large group of false-positives, with respect to cochlear pathology, that need to be identified. This is because the next stage of the process is to evaluate all positive cases using a much more time consuming and costly procedure, such as behavioral testing and/or Auditory Brainstem Evoked Response Audiometry (ABER).
The use of evoked otoacoustic emissions as a screening tool is growing rapidly because of the speed, ease of testing and objectivity of this technique. In order for a screening program using evoked otoacoustic emissions to be cost effective it is essential that the high rate of false positives with respect to cochlear pathology due to middle-ear problems be reduced substantially. An effective solution to this problem would be to separate cochlear "positives" from middle-ear "positives" during the evoked otoacoustic emissions screening stage and to initiate appropriate forms of evaluation and intervention for each of these cases.
All of the problems described above can either be totally eliminated, or at least reduced considerably, if we knew the acoustic power flow in the ear canal. In addition to these advantages there are other, more subtle considerations for developing an instrument capable of measuring acoustic power flow in the ear canal.
In the normal human middle ear, the ear canal and middle-ear impedances are substantially matched for frequencies above 800 Hz, allowing for efficient power flow from the ear canal to the cochlea. Ear drum impedance will change when the middle-ear malfunctions, for example due to static pressure in the middle-ear space, or under more serious conditions such as changes in the ossicular impedance due to otospongiosis (otosclerosis), or changes in the stiffness of the ossicular ligaments. When impedance mismatch is large, power reflectance approaches 1 (i.e., 100% of the power is reflected). Under this condition output acoustic intensity is nearly equal to the input acoustic intensity, and the forward and retrograde waves are almost equal in magnitude. As a result, the pressure nearly cancels at frequencies corresponding to 1/4 acoustic wavelengths or 1/2 wavelength round trips from the reflection point. Pressure will be very small at the measurement point for this frequency yielding a misleading pressure calibration since the ear canal pressure at the measurement point for this frequency is not a useful or accurate measure of either drum pressure or the power absorbed by the middle-ear and cochlea.
Measuring the power absorbed by the middle ear and cochlea provides many advantages in audiology, particularly when at frequencies above 4 kHz where there is a middle-ear pathology, or when the physical size of the ear is very different from the normal adult ear. This is typical in cases where traditional coupler measurements provide misleading estimates of the sound pressure level at the eardrum. Three important observations have been derived for normal ears. First, middle-ear impedance mismatch is a large source of variability from human to human. Second, a cat's middle ear is nearly a lossless system. Third, power flux at the threshold of hearing in a gerbil has been found to be constant.
The first observation identifies a significant physical source of calibration variability. The second observation means that virtually all of the power flowing into the middle ear is delivered to the cochlea. The third observation is consistent with the idea that acoustic power flow is an important correlate of cochlear hearing thresholds.
In order to measure power absorbed by the middle ear it is necessary to measure the acoustic impedance of the middle-ear. Instruments in current clinical use for measuring acoustic impedance of the ear, however, do not measure acoustic impedance directly, but rather measure the relative impedance magnitude, i.e., the impedance magnitude relative to that for a normal ear. These instruments are also frequently limited to a few standard test frequencies (e.g., 220 Hz and 600 Hz), rather than providing a measurement of acoustic impedance over the entire audio frequency range.
Recent advances in transducer development and concomitant advances in computerized measurement of sound transmission characteristics in the ear allow for a practical means of measuring acoustic impedance, and more importantly, sound power absorption by the ear. The instrument of the present invention is a significant improvement over a technique developed by Jont B. Allen, "Measurement of Eardrum Acoustic Impedance", Peripheral Auditory Mechanism, pp. 44-51 (1986). In Allen's methodology two transducers are inserted into the ear canal. One transducer generates a test sound, the other, a microphone, measures the pressure in the test ear. The two transducers are then placed in four cavities, where the cavity pressures are measured. This information is then analyzed to produce an estimate of the canal reflectance. The above described measurement method has been applied to animals, adults, temporal bones, and infants. The instrumentation used in Allen's experiments, however, is expensive and complex and is therefore not amenable to use in a clinical setting.